A normal ear transmits sounds as shown in FIG. 1 through the outer ear 101 to the tympanic membrane 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, auditory prostheses have been developed. For example, when the impairment is related to operation of the middle ear 103, a conventional hearing aid may be used to provide acoustic-mechanical stimulation to the auditory system in the form of amplified sound. Or when the impairment is associated with the cochlea 104, a cochlear implant with an implanted stimulation electrode can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
FIG. 1 also shows some components of a typical cochlear implant system which includes an external microphone that provides an audio signal input to an external signal processor 111 where various signal processing schemes can be implemented. The processed signal is then converted into a digital data format, such as a sequence of data frames, for transmission into the implant 108. Besides receiving the processed audio information, the implant 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces a stimulation pattern (based on the extracted audio information) that is sent through an electrode lead 109 to an implanted electrode array 110.
Typically, this electrode array 110 includes multiple stimulation contacts 112 on its surface that provide selective stimulation of the cochlea 104. Depending on context, the stimulation contacts 112 are also referred to as electrode channels. In cochlear implants today, a relatively small number of electrode channels are each associated with relatively broad frequency bands, with each stimulation contact 112 addressing a group of neurons through an electric stimulation pulse having a charge which is derived from the instantaneous amplitude of the signal envelope within that frequency band.
In some coding strategies, stimulation pulses are applied at a constant rate across all electrode channels, whereas in other coding strategies, stimulation pulses are applied at a channel-specific rate. Various specific signal processing schemes can be implemented to produce the electrical stimulation signals. Signal processing approaches that are well-known in the field of cochlear implants include continuous interleaved sampling (CIS) digital signal processing, channel specific sampling sequences (CSSS) digital signal processing (as described in U.S. Pat. No. 6,348,070, incorporated herein by reference), spectral peak (SPEAK) digital signal processing, and compressed analog (CA) signal processing.
For example, FIG. 2 shows the major functional blocks in a typical cochlear implant signal processing system wherein band pass signals are processed and coding to generate electrode stimulation signals to stimulation electrodes in an implanted cochlear implant electrode array. Preprocessor Filter Bank 201 pre-processes the initial acoustic audio signal with a bank of band pass filters, each of which is associated with a specific band of audio frequencies so that the acoustic audio signal is filtered into some M band pass signals, B1 to BM where each signal corresponds to the band of frequencies for one of the band pass filters. Based on the tonotopic organization of the cochlea, each stimulation contact in the scala tympani often is associated with a specific band pass filter of the external filter bank. FIG. 3 shows an example of a short time period of an audio speech signal from a microphone, and FIG. 4 shows an acoustic microphone signal decomposed by band-pass filtering by a bank of filters into a set of signals.
The band pass signals B1 to BM are input to a Signal Processor 202 which extracts signal specific stimulation information—e.g., envelope information, phase information, timing of requested stimulation events, etc.—into a set of N stimulation channel signals S1 to SN that collectively represent the sound information that is present in the initial acoustic audio signal. Stimulation Coding Module 203 then converts the processed stimulation channel signals S1 to SN to produce a corresponding sequence of electrode stimulation signals A1 to AM that provide an optimal electric representation of the acoustic signal, while the Pulse Mapping and Shaping Module 204 then applies a linear mapping function (typically logarithmic) and pulse shaping of the electrode stimulation signals A1 to AM that is adapted to the needs of the individual implant user based on a post-surgical fitting process that determines patient-specific perceptual characteristics.
The output of the Pulse Mapping and Shaping Module 204 is a set of electrode stimulation signals E1 to EM to the stimulation contacts in the implanted electrode array which stimulate the adjacent nerve tissue. Symmetrical biphasic current pulses are often applied for stimulation. In the specific case of a CIS system, the stimulation pulses are applied in a strictly non-overlapping sequence. Thus, as a typical CIS-feature, only one electrode channel is active at one time and the overall stimulation rate is comparatively high. For example, assuming an overall stimulation rate of 18 kpps and a 12 channel filter bank, the stimulation rate per channel is 1.5 kpps. Such a stimulation rate per channel usually is sufficient for adequate temporal representation of the envelope signal. The maximum overall stimulation rate is limited by the minimum phase duration per pulse. The phase duration cannot be chosen arbitrarily short, because the shorter the pulses, the higher the current amplitudes have to be to elicit action potentials in neurons, and current amplitudes are limited for various practical reasons. For an overall stimulation rate of 18 kpps, the phase duration is 27 μs, which is near the lower limit. Each output of the CIS band pass filters can roughly be regarded as a sinusoid at the center frequency of the band pass filter which is modulated by the envelope signal. This is due to the quality factor (Q≈3) of the filters. In case of a voiced speech segment, this envelope is approximately periodic, and the repetition rate is equal to the pitch frequency.
In the existing CIS-strategy, only the signal envelopes are used for further processing, i.e., they contain the entire stimulation information. For each electrode channel, the signal envelope is represented as a sequence of biphasic pulses at a constant repetition rate. A characteristic feature of CIS is that this repetition rate (typically 1.5 kpps) is equal for all electrode channels and there is no relation to the center frequencies of the individual channels. It is intended that the repetition rate is not a temporal cue for the patient, i.e., it should be sufficiently high, so that the patient does not perceive tones with a frequency equal to the repetition rate. The repetition rate is usually chosen at greater than twice the bandwidth of the envelope signals (Nyquist theorem).
Another cochlear implant stimulation strategy that transmits fine time structure information is the Fine Structure Processing (FSP) strategy by Med-El. Zero crossings of the band pass filtered time signals are tracked, and at each negative to positive zero crossing a Channel Specific Sampling Sequence (CSSS) is started. Typically CSSS sequences are only applied on the first one or two most apical electrode channels, covering the frequency range up to 200 or 330 Hz. The FSP arrangement is described further in Hochmair I, Nopp P, Jolly C, Schmidt M, SchoBer H, Garnham C, Anderson I, MED-EL Cochlear Implants: State of the Art and a Glimpse into the Future, Trends in Amplification, vol. 10, 201-219, 2006, which is incorporated herein by reference.
Many CI coding strategies use what is referred to as an N-of-M approach where only some number n electrode channels with the greatest amplitude are stimulated in a given sampling time frame. If, for a given time frame, the amplitude of a specific electrode channel remains higher than the amplitudes of other channels, then that channel will be selected for the whole time frame. Subsequently, the number of electrode channels that are available for coding information is reduced by one, which results in a clustering of stimulation pulses. Thus, fewer electrode channels are available for coding important temporal and spectral properties of the sound signal such as speech onset.
One method to reduce the spectral clustering of stimulation per time frame is the MP3000™ coding strategy by Cochlear Ltd, which uses a spectral masking model on the electrode channels. Another method that inherently enhances coding of speech onsets is the ClearVoice™ coding strategy used by Advanced Bionics Corp, which selects electrode channels having a high signal to noise ratio. U.S. Patent Publication 2005/0203589 describes how to organize electrode channels into two or more groups per time frame. The decision which electrode channels to select is based on the amplitude of the signal envelopes.
In addition to the specific processing and coding approaches discussed above, different specific pulse stimulation modes are possible to deliver the stimulation pulses with specific stimulation contacts—i.e. mono-polar, bi-polar, tri-polar, multi-polar, and phased-array stimulation. And there also are different stimulation pulse shapes—i.e. biphasic, symmetric triphasic, asymmetric triphasic pulses, or asymmetric pulse shapes. These various pulse stimulation modes and pulse shapes each provide different benefits; for example, higher tonotopic selectivity, smaller electrical thresholds, higher electric dynamic range, less unwanted side-effects such as facial nerve stimulation, etc. But some stimulation arrangements are quite power consuming, especially when neighboring stimulation contacts are used as current sinks. Up to 10 dB more charge might be required than with simple mono-polar stimulation concepts (if the power-consuming pulse shapes or stimulation modes are used continuously).
Another consideration as to stimulation pattern is the spread of the excitation pattern in the stimulated neural tissue. This excitation spread is a function of amplitude so that at relatively low stimulation levels near the hearing threshold level, the region of excited neurons is small, whereas at high stimulation levels, large populations throughout the cochlea are excited. See, e.g., U.S. Pat. No. 7,941,223, which is incorporated herein by reference in its entirety. Excitation spread occurs both with acoustic excitation (as in normal hearing) and with electric stimulation of the neural tissue as in a cochlear implant.
Some literature in the field discusses stimulation modes intended to produce selective stimulation; e.g. U.S. Pat. No. 7,899,547; EP 2482923, and Litvak et al., Loudness growth observed under partially tripolar stimulation: Model and data from cochlear implant listeners, J. Acoust. Soc. Am. 122 2, August 2007; all of which are incorporated herein by reference in their entireties. Such literature takes into account the natural behavior of a level-dependent spread of excitation. At high perceptual levels, stimulation modes intending to generate focused stimulation, produce a spread of excitation that is quite similar to the spread of a simple mono-polar stimulation mode. For a wide spread of excitation, mono-polar stimulation is probably the most power-efficient mode, whereas at low stimulation levels, multi-polar stimulation is known to achieve the lowest spread of excitation, albeit in a less power efficient manner.